The field of the invention is imaging methods and systems. More particularly, the invention relates to a method for improving the resolution of projected images.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, M.sub.z, may be rotated, or "tipped", into the x-y plane to produce a net transverse magnetic moment M.sub.t. A signal is emitted by the excited spins, and after the excitation signal B.sub.1 is terminated, this signal may be received and processed to form an image.
When utilizing these signals to produce images, magnetic field gradients (G.sub.x, G.sub.y and G.sub.z) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The present invention will be described with reference to the well known Fourier transform (FT) imaging technique. This technique is discussed, for example in an article entitled "Spin Warp NMR Imaging and Applications to Human Whole-Body Imaging" by W. A. Edelstein et al., Physics in Medicine and Biology, Vol. 25, pp. 751-756 (1980). It employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of NMR spin-echo or gradient-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (2DFT), for example, spatial information is encoded in one direction by applying a phase encoding gradient (G.sub.y) along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient (G.sub.x) in a direction orthogonal to the phase encoding direction. The readout gradient present during the echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse G.sub.y is incremented (.DELTA.G.sub.y) in the sequence of views that are acquired during the scan and the NMR signal is sampled in the presence of the readout gradient G.sub.x to produce a set of NMR data from which an entire image can be reconstructed. The acquired two-dimensional array of "k-space" NMR data is Fourier transformed to produce a corresponding image. The same method can be expanded to produce three-dimensional images by stepping a third gradient (G.sub.z) through a sequence of phase encoding values during the scan.
The resolution of the reconstructed image is determined in large part by the number of signal samples acquired during the scan. For example, if the phase encoding gradient is stepped through 256 levels and if 256 samples of each NMR signal are acquired during the scan the resulting k-space NMR data array will contain 256 by 256 data elements. When this is Fourier transformed, an image array containing 256 by 256 pixels is reconstructed. On the other hand, if the number of phase encodings is reduced to 128, the resolution of the resulting 128 by 256 pixel image is reduced along one axis. Since it requires more scan time to acquire more samples, most clinical scans are a compromise between the need for a higher resolution image and the need for a reduced scan time.
Magnetic resonance angiography ("MRA") uses the nuclear magnetic resonance phenomenon to produce images of the human vasculature. Two basic techniques have been proposed and evaluated. The first class, time-of-flight (TOF) techniques, consists of methods which use the motion of the blood relative to the surrounding tissue. The most common approach is to exploit the differences in signal saturation that exist between flowing blood and stationary tissue. Flowing blood, which is moving through the excited section, is continually refreshed by spins experiencing fewer excitation pulses and is, therefore, less saturated. The result is the desired image contrast between the high-signal blood and the low-signal stationary tissues.
MRA methods have also been developed that encode motion into the phase of the acquired signal as disclosed in U.S. Pat. No. Re. 32,701. These form the second class of MRA techniques and are known as phase contrast (PC) methods. Currently, most PC MRA techniques acquire two images, with each image having a different sensitivity to the same velocity component. Angiographic images are then obtained by forming either the phase or complex difference between the pair of velocity-encoded images.
Despite the tremendous strides made in recent years, at many clinical sites MRA is still considered a research tool and is not routinely used in clinical practice. More widespread application of either TOF or PC techniques is hampered by the presence of a variety of deleterious image artifacts, which can mask and, in some cases, even mimic pathology. One of these deleterious effects is lack of clear vessel edge definition, particularly with small vessels. Compromised edge definition may be due to partial volume averaging or due to the position of the vessel edge relative to the reconstructed array of voxels.
There are two methods used to improve MRA vessel edge definition when increased sampling is not an available option to increase resolution. The first technique is to reposition the image reconstruction grid so that the vessel edge lies entirely within a reconstructed voxel. This is easily accomplished by imparting a uniform phase shift to all of the acquired k-space data. The difficulty with this solution is that it may improve the definition of some vessel edges and reduce the definition of others.
The second technique is to interpolate the reconstructed image data to a finer grid spacing. Linear interpolation and cubic spline interpolation can be used, for example, to improve image quality, but such real-space interpolation techniques do not increase the resolution of the image.
Another method which is used to improve image resolution is referred to in the art as "zero-filled interpolation" or "sinc interpolation" or "band-limited interpolation". As described by Y. P. Du et al, "Reduction of Partial-Volume Artifacts with Zero-Filled Interpolation in Three-Dimensional MR Angiography," JMRI 1994; 4:733-741, zero-filled interpolation is usually implemented by appending zeros on each dimension of the k-space data before it is Fourier transformed. In 3D MR angiography, for example, zeros are appended on each dimension of k-space in all three spatial frequency directions. The much enlarged array of k-space data is then inverse Fourier transformed to produce a corresponding enlarged image data array. The field of view is not changed in this enlarged image array, but the number of pixels required to depict a particular structure is increased. Substantial improvement in vessel continuity and visibility, especially in small vessels, is achieved using this method.
Zero-filled interpolation is difficult to implement in commercially available MRI systems. The k-space data set is increased in size by a factor of two or more along each of its dimensions. For a 3D angiographic data set, this means an increase in the k-space data storage and the image storage by a factor of eight or more. This makes it more costly and difficult to archive, network and display. In addition, the increased k-space data array requires much more time to Fourier transform with existing array processor hardware.
Yet another approach described by T. O. Cooper et al "Improved Magnetic Resonance Angiography Vessel Edge Definition With Truncated Sinc Interpolation" ISMRM Proceedings, Vol. 3, April/May 1996. A truncated sinc function is proposed for interpolation in the spatial domain. To keep the computational burden within reason, this method compromises interpolation accuracy.